Embodiments of the invention relate generally to diagnostic imaging and, more particularly, to an apparatus and method of detecting x-rays.
Typically, in x-ray systems, such as computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped or cone-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces an electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis, which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about a gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors also typically include an anti-scatter grid (sometimes called post-patient collimator) for eliminating scattered x-rays arriving at the detector, a scintillator for converting x-rays to light energy, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom. Typically, each scintillator element of a scintillator array converts x-rays to light energy, which is optically guided to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
Current CT detectors generally use detectors such as scintillation crystal/photodiode arrays, where the scintillation crystal absorbs x-rays and converts the absorbed energy into visible light. These arrays are often based on front-illuminated photodiodes. However, for products where the number of detector rows is beyond 64, the designs are generally based on back-illuminated photodiodes.
A development of multi-slice CT systems has led the market to new applications in general and to cardiac and perfusion imaging in particular. A goal and/or desire of many clinicians is to image a heart within one gantry rotation (or within a half-scan or half a gantry rotation) and with improved temporal resolution. To address such goals or desires, detectors having a large coverage (e.g., system coverage up to 160 mm or more at iso-center) have been investigated and developed. Such detectors generally have a large number of detector rows (e.g., 256 rows or more) and are correspondingly capable of acquiring data corresponding to a large number of slices (e.g., 256 slices or more) during one scan or gantry rotation. It is noted, however, that detector costs generally increase as its number of detector rows increase.
Not all applications, however, greatly benefit from high-detector-row-count acquisitions such as 256 or more detector-row-acquisitions. For example, many conventional types of CT imaging do not require the increased coverage obtained by the use of 256-row detectors. As such, using a large row-count detector in many instances can be “overkill.” To address this situation, technicians often employ more than one type of CT scanner. For example, when large coverage is needed, a technician may employ a 256-detector-row CT scanner, and when large coverage is not need, a technician may employ a 64-detector-row CT scanner. The use of multiple types of CT scanners, however, can be cost prohibitive because of the costs associated with purchasing multiple CT scanners.
It would therefore be beneficial to design a cost effective system including an x-ray detector capable of varying detector-row number.